Optical imaging is an evolving clinical imaging modality that uses penetrating lights rays to create images of both intrinsic and extrinsic biological scatterers. Light offers unique contrast mechanisms that can be based on absorption, e.g., probing of hemoglobin concentration or blood saturation, and/or fluorescence, e.g., probing for weak auto-fluorescence, or exogenously administered fluorescent probes (Neri et al., Nat. Biotech. 15:1271-1275, 1997; Ballou et al., Cancer Immunol. Immunother. 41:257-63,1995; and Weissleder, et al., Nat. Biotech. 17:375-178, 1999). Preferably, light in the red and near infrared range (600-1200 nm) is used to maximize tissue penetration and minimize absorption from natural biological absorbers such as hemoglobin and water. (Wyatt, Phil. Trans. R. Soc. London B 352:701-706, 1997; Tromberg, et al., Phil. Trans. R. Soc. London B 352:661-667, 1997).
Diffuse optical tomography (DOT) is one type of optical tomography that has been used for quantitative, three-dimensional imaging of biological tissue, based on intrinsic absorption and scattering. (Ntziachristos et al., Proc. Natl. Acad. Sci. USA, 97:2767-72, 2000; Benaron et al., J. Cerebral Blood Flow Metabol. 20:469-77, 2000) A typical DOT imaging system uses narrow-band light sources so that specific extrinsic and/or intrinsic fluorophores are targeted. Light, customarily generated by laser diodes, is usually directed to and from the target tissue using fiber guides, since (1) it enables flexibility in the geometrical set-up used, (2) reduces signal loss, and (3) simplifies the theoretical modeling of contact measurements. The use of fiber guides, however, has significant disadvantages. The most significant is that only a limited number of detector channels can be implemented since scaling up requires a large number of fibers (usually fiber bundles) that have to be coupled to the tissue, which in many cases is not practical. In addition, it is also very difficult to control or measure the exact coupling conditions of each individual fiber or fiber bundle, which can vary quite significantly from fiber to fiber. An alternative to fibers is to use a compression plate and/or an optical matching fluid. For example, it is common to compress the tissue of investigation into a fixed geometry such as a slab or circle and to use an optical matching fluid to eliminate possible air-tissue interfaces. In either case, the use of fiber guides and/or optical matching fluids and fixed geometries impedes the experimental practicality and severely limits the tomographic capacity of the imaging system. These constraints result in significant limitations to the use of these systems either in research or in the clinic. To date, non-contact systems and methods for tomography of biological tissue and other diffuse and diffuse-like medium with arbitrary boundaries have not been developed or reported.
Currently, optical tomography utilizes either numerical or analytical methods for solving the equations governing propagation of light through biological tissue. Numerical methods, such as the finite element method (FEM), finite differences (FD) or the boundary element method (BEM), are used for complex geometries of air/tissue boundaries and are extremely computationally costly and therefore are currently non-viable in a real-time three-dimensional research and clinical setting. (For example, reported numerical-based reconstruction times for a simulated typical 3D breast-imaging problem on state of the art single processor computers range from 6 to 36 hours.) Analytical methods are much faster (for example, an analytical-based 3D reconstruction case of the breast ranges between 2-15 minutes), but are available only for simple geometries of air/tissue boundary such as a slab, a cylinder or a sphere, and often lack adequate accuracy for imaging complex objects such as a human breast.
An alternative method to perform tomography of diffuse or diffuse-like medium with complex geometries is by means of an analytical approach called the Kirchhoff Approximation (KA). This method uses the angular spectrum representation of the propagating average intensity and employs the reflection coefficients for light waves to calculate the light intensity inside any arbitrary geometry. Although in contact imaging applications the KA can achieve relatively good computational efficiency (Ripoll et al., Opt. Lett. 27:333-335, 2002), it has several significant limitations that would restrict its use in real research and clinical settings. Specifically, the KA is limited to geometries such as a cylinder or ellipse, i.e. geometries that do not include shadow regions. Furthermore, the KA method generally works for larger volumes (e.g., diameters >3 cm), and for highly absorbing medium (e.g., typically the absorption coefficient must be 10-100 times higher than that of water).